Optical coherence tomography (OCT) is a novel biomedical imaging technique that can render 2D and 3D structural and functional information in real time.1,2 OCT is based on the theory of low-coherence interferometry. Biological samples absorb very little and scatter some of the near infrared light (NIR) that they are probed with.2 OCT uses NIR to probe specimens as deep as a few millimeters, with micron resolution. OCT systems have been introduced recently in a clinical setting for use in ophthalmology.
In OCT the NIR probing light is equally split into a mirror arm that serves as a reference and a sample arm. The interference of the backscattered light fields in the two arms of the interferometer (Er and Es) is detected (as intensity Id) and used to determine the structure of the scattering object in the sample arm. Image reconstruction is based on the theory of inverse scattering; by inverse Fourier-transforming the autocorrelation signal from the demodulated detected intensity at different times (time-domain OCT, or TD-OCT, FIG. 1) or wavelengths (spectral-domain OCT, or SD-OCT; also commonly referred to as Fourier domain OCT), one can retrieve the complex analytic signal that contains amplitude and phase information about the object. For interference to occur in TD-OCT the optical paths of the reference and the sample arms need to differ by no more than the coherence length Ic of the source. This also sets the limit on the axial resolution of the system, which is equal to Ic . For a Gaussian probing spectrum, Ic is inversely proportional to the bandwidth Δλ of the source. Therefore, the use of a very broad bandwidth source for high axial resolution imaging is desired. The transverse resolution RT is given by the diameter of the probing beam 2w0 and can be expressed in terms of the focal length f of the collimator, the center wavelength of the source λ0, and the diameter of the focused beam D, as shown below (assuming a Gaussian probing beam).
                                          I            d                    ⁡                      (            t            )                          =                ⁢                                                            (                                                      I                    r                                    +                                      I                    s                                                  )                            2                        +            Re                    <                                                    E                r                *                            ⁡                              (                                  t                  +                  τ                                )                                      ⁢                                          E                s                            ⁡                              (                t                )                                              >                                                  l          c                =                ⁢                                                            2                ⁢                                                                  ⁢                ln                ⁢                                                                  ⁢                2                            π                        ⁢                                          λ                0                2                                            Δ                ⁢                                                                  ⁢                λ                                              ≈                      0.44            ⁢                                          λ                0                2                                            Δ                ⁢                                                                  ⁢                λ                                                                                      R          T                =                ⁢                              2            ⁢                                                  ⁢                          w              0                                ≈                      2.44            ⁢                                          f                ⁢                                                                  ⁢                                  λ                  0                                            D                                          
Superparamagnetic iron oxide (SPIO) particles have been used extensively as contrast agents for magnetic resonance imaging (MRI).9 Magnetic particles with small core sizes (<100 nm) are easily transported through the circulatory system and are able to extravasate, and are thus suitable for both in vivo and in vitro studies.6,8 Depending on their composition and size, magnetic particles can be very responsive to external, non-invasive manipulation or detection due to their strong magnetic susceptibility. Moreover, they can be functionalized to target antigens and thus enhance contrast at the molecular and cellular level, aiding in pathogen localization and early diagnosis of disease. The use of these magnetic particles in OCT has several advantages: the ability to externally manipulate the particles, the low magnetic susceptibility inherent in human tissues, the availability of FDA approved biocompatible iron oxide particles for MRI contrast, and the potential for hyperthermic therapy with high frequency (>100 kHz) modulation.
Magnetomotive optical coherence tomography (MM-OCT) in a time-domain optical coherence tomography (TD-MMOCT) system has been used for detecting the displacements in different samples caused by the modulation of the magnetic field and it has been subsequently shown that the magnetomotive response in the system is predictable.8 In this scheme, axial scans in a two-dimensional transversal sample plane are acquired with the magnetic field off and on, while allowing the particles and the sample sufficient time to complete motion and reach equilibrium between axial scans, for example at a line rate of 10 Hz. Thus, the images taken with the TD-MMOCT system represent a static description of the sample in the absence and in the presence of the magnetic field, and may be used as a background-rejecting method by estimating a background displacement signal when the magnetic field is off, compared to the magnetic-specific displacement when the magnetic field is off-on.8 
This previous work demonstrated the ability to image magnetite (Fe3O4) micro-and nanoparticles after uptake by in vitro macrophages4 and in vivo African frog tadpoles8 by modulating an externally applied magnetic field and detecting the resultant magnetomotion specific to the particles. Other researchers have also used this principle to provide hemoglobin contrast in optical Doppler tomography,31 and to detect iron uptake in tissues with differential phase OCT32 and also in ultrasound.9 
Phase measurements in common-path low-coherence light interferometry have been shown to render high sensitivity to sub-wavelength displacements or obstacles in the path of light.10-12 Path length sensitivities as low as 25 m for spectral-domain optical coherence phase microscopy (SD-OCPM)10 and 18 m (equivalent phase stability=0.4 mrad) for spectral-domain phase microscopy (SDPM)11 have been reported. Phase-resolved methods10-15 are often used in a dynamical regime, such as in measuring intralipid16-18 or blood flow19-23 velocities, nerve displacements,24 or monitoring cell10 and even cardiomyocyte12 activity.